Direct conversion energy discriminating CT detector with over-ranging correction

ABSTRACT

A CT detector capable of energy discrimination and direct conversion is disclosed. The detector includes multiple layers of semiconductor material with the layers having varying thicknesses. The detector is constructed to be segmented in the x-ray penetration direction so as to optimize count rate performance as well as avoid saturation. The detector also includes variable pixel pitch and a flexible binning of pixels to further enhance count rate performance.

CROSS-REFERENCE TO RELATED APPLICATION

The present application is a continuation of and claims priority of U.S.patent application Ser. No. 10/939,787 filed Sep. 13, 2004, thedisclosure of which is incorporated herein.

BACKGROUND OF THE INVENTION

The present invention relates generally to radiographic detectors fordiagnostic imaging and, more particularly, to a multi-layer directconversion CT detector capable of providing photon count and/or energydata with improved saturation characteristics and over-rangingself-correctability.

Typically, in radiographic imaging systems, such as x-ray and computedtomography (CT), an x-ray source emits x-rays toward a subject orobject, such as a patient or a piece of luggage. Hereinafter, the terms“subject” and “object” may be interchangeably used to describe anythingcapable of being imaged. The beam, after being attenuated by thesubject, impinges upon an array of radiation detectors. The intensity ofthe attenuated beam radiation received at the detector array istypically dependent upon the attenuation of the x-rays. Each detectorelement of the detector array produces a separate electrical signalindicative of the attenuated beam received by each detector element. Theelectrical signals are transmitted to a data processing system foranalysis which ultimately produces an image.

In other typical radiographic imaging systems, positron emissiontomography (PET) or single photon emission computed tomography (SPECT) aradiation source within the imaged object emits x-rays which areintercepted by a photon counting, energy sensitive x-ray detector. A CTsystem can be paired with a PET or SPECT system to produce a fusedsystem (CT/SPECT or CT/PET) providing images indicating both anatomicalstructure and physiologically significant (i.e. functional) information.Such combined systems include a source that emits x-rays toward a x-raydetector and separate SPECT or PET detector which measures x-raysemitted from radiation source within the object.

In some CT imaging systems, for example, the x-ray source and thedetector array are rotated within a gantry and within an imaging planearound the subject. X-ray sources for such CT imaging systems typicallyinclude x-ray tubes, which emit the x-rays as a fan beam emanating froma focal point. X-ray detectors for such CT imaging systems typically areconfigured in an circular arc centered to the focal spot. In addition,such detectors include a collimator for collimating x-ray beams receivedat the detector with focus to the focal spot. In addition, suchdetectors include a scintillator for converting x-rays to light energyadjacent the collimator, and a photodiode for receiving the light energyfrom an adjacent scintillator and producing electrical signalstherefrom. Typically, each scintillator of a scintillator array convertsx-rays to light energy. Each photodiode detects the light energy andgenerates a corresponding electrical signal as a function of the lightemitted by a corresponding photodiode. The outputs of the photodiodesare then transmitted to the data processing system for imagereconstruction.

In some SPECT and PET systems, for example, the one or more flatdetector arrays is rotated within a gantry and within an imaging planeand around the subject. X-ray radiation sources within the imaged objectemit photons in random directions. A x-ray detector typically includes acollimator for collimating x-ray beams received at the detector withfocus for parallel rays contained within the imaging plane andperpendicular to the detector plane. In addition, such detectors includea scintillator for converting x-rays to light energy adjacent thecollimator, and a photomultiplier tube for receiving the light energyfrom an adjacent scintillator and producing electrical signals therefromwhich are then transmitted to the data processing system for imagereconstruction.

Conventional CT imaging systems utilize detectors that convertradiographic energy into current signals that are integrated over a timeperiod, then measured and ultimately digitized. A drawback of suchdetectors however is their inability to provide data or feedback as tothe number and/or energy of photons detected. That is, conventional CTdetectors have a scintillator component and photodiode component whereinthe scintillator component illuminates upon reception of radiographicenergy and the photodiode detects illumination of the scintillatorcomponent and provides an electrical signal as a function of theintensity of illumination. While it is generally recognized that CTimaging would not be a viable diagnostic imaging tool without theadvancements achieved with conventional CT detector design, a drawbackof these detectors is their inability to provide energy discriminatorydata or otherwise count the number and/or measure the energy of photonsactually received by a given detector element or pixel. That is, thelight emitted by the scintillator is a function of the number of x-raysimpinged as well as the energy level of the x-rays. Under the chargeintegration operation mode, the photodiode is not capable ofdiscriminating between the energy level or the photon count from thescintillation. For example, two scintillators may illuminate withequivalent intensity and, as such, provide equivalent output to theirrespective photodiodes. Yet, the number of x-rays received by eachscintillator may be different as well as the x-rays intensity, but yieldan equivalent light output.

A typical PET or SPECT system uses a photon counting, energydiscriminating detector constructed from a scintillator andphotomultiplier tube. Such detectors have large detector elements and assuch are not readily adapted to CT applications requiring highresolution imaging to capture anatomical detail in the imaged object.Accordingly, recent detector developments have included the design of anenergy discriminating, direct conversion detector that can providephoton counting and/or energy discriminating feedback with high spatialresolution. In this regard, the detector can be caused to operate in anx-ray counting mode, an energy measurement mode of each x-ray event, orboth.

These energy discriminating, direct conversion detectors are capable ofnot only x-ray counting, but also providing a measurement of the energylevel of each x-ray detected. Consequently, such a detector couldpotentially be used for SPECT or PET imaging. While a number ofmaterials may be used in the construction of a direct conversion energydiscriminating detector, semiconductors have been shown to be onepreferred material.

A drawback of direct conversion semiconductor detectors, however, isthat these types of detectors cannot count at the x-ray photon fluxrates typically encountered with conventional CT systems, e.g. at orabove 1 million counts per sec per millimeter squared (1.0 Mcps). Thevery high x-ray photon flux rate, above 1.0 Mcps, causes pile-up andpolarization which ultimately leads to detector saturation. That is,these detectors typically saturate at relatively low x-ray flux levelthresholds. Above these thresholds, the detector response is notpredictable or has degraded dose utilization. For SPECT and PET, imagingflux levels are below 1.0 Mcps and such saturation in a semiconductordetector for SPECT and PET is not a practical concern. However, for CT,saturation can occur at detector locations wherein small subjectthickness is interposed between the detector and the radiographic energysource or x-ray tube. It has been shown that these saturated regionscorrespond to paths of low subject thickness near or outside the widthof the subject projected onto the detector fan-arc. In many instances,the subject is more or less circular or elliptical in the effect onattenuation of the x-ray flux and subsequent incident intensity to thedetector. In this case, the saturated regions represent two disjointedregions at extremes of the fan-arc. In other less typical, but not rareinstances, saturation occurs at other locations and in more than twodisjointed regions of the detector. In the case of an ellipticalsubject, the saturation at the edges of the fan-arc is reduced by theimposition of a bowtie filter between the subject and the x-ray source.Typically, the filter is constructed to match the shape of the subjectin such a way as to equalize total attenuation, filter and subject,across the fan-arc. The flux incident to the detector is then relativelyuniform across the fan-arc and does not result in saturation. What canbe problematic, however, is that the bowtie filter may not be optimalgiven that a subject population is significantly less than uniform andnot exactly elliptical in shape. In such cases, it is possible for oneor more disjointed regions of saturation to occur or conversely toover-filter the x-ray flux and create regions of very low flux. Lowx-ray flux in the projection will ultimately contribute to noise in thereconstructed image of the subject.

Detector saturation causes loss of imaging information and results inartifacts in x-ray projection and CT images. In addition, hysteresis andother non-linear effects occur at flux levels near detector saturationas well as flux levels over detector saturation. Direct conversiondetectors are susceptible to a phenomenon called “polarization” wherecharge trapping inside the material changes the internal electric field,alters the detector count and energy response in an unpredictable way,and results in hysteresis where response is altered by previous exposurehistory. In particular, photon counting, direct conversion detectors,saturate due to the intrinsic charge collection time (i.e. dead time)associated with each x-ray photon event. Saturation will occur due topulse pile-up when x-ray photon absorption rate for each pixel is on theorder of the inverse of this charge collection time. The chargecollection time is approximately proportional to thickness of the directconversion layer for a fixed electric field and anode contact size;therefore, an increase in saturation rate is possible if the directconversion layer is thinner. However, a sufficient thickness is requiredto stop almost all the x-rays. Incomplete collection of x-rays resultsin reduced image quality, i.e. a noisy image, and poor utilization ofdose to the imaged object.

An additional factor in the charge collection time is the voltageapplied across the layer thickness. A larger electric field(voltage/thickness) results in inverse proportionally smaller chargecollection times and proportionally larger saturate rates. However,there is a reliability issue to routing of high voltage signals. Higherreliability can be obtained for lower voltages across smallerthicknesses of direct conversion layer. However, again, a sufficientthickness of the layer is required to sufficiently stop a majority ofthe x-rays.

Other types of detectors in addition to direct conversion detectors alsosaturate. A common example is the scintillator-photodiode arrangementconnected to an integrating preamplifier. Charge created from eachphoton is routed to the preamplifier. As x-ray flux increases, thecurrent to the preamplifier or total charge built up over an integrationtime period will increase. The readout electronics have a limitingcurrent or charge capability before saturating the amplifier. Amplifiersaturation is associated with non-linear response and the loss of signalcharge. This again results in poor dose utilization and image artifacts.

Another detector construction is a scintillator over photodiodeconnected to photon counting readout electronics. Similar constructionsutilize a scintillator over avalanche-photodiode or photo-multipliertube. Saturation of the x-ray flux rate in these photon-counting casesis also related to a dead time for clearing the charge before arrival ofthe next x-ray photon.

For photon counting, direct conversion detectors, a practical solutionto x-ray flux rate saturation in imaging systems using x-ray sourcesoperating at or above 1.0 Mcps range is not known. For these systems, atotal thickness of the x-ray absorbing layer must be greater than 1.0mm. The higher the energy of the x-rays; the higher the requiredthickness to sufficiently stop a predominance of the x-ray flux. Atypical target value is to stop 95% or greater of the incident x-rays.For Cadmium Zinc Telluride (CZT) or Cadmium Telluride (CdTe), twopossible direct conversion materials used for x-ray spectroscopy, therequired thickness for diagnostic radiology and CT imaging is 3.0-5.0 mmin order to stop most of the x-rays generated from a source at 100-200kVp. For CZT and CdTe, the saturation limit of 10⁷ x-rays/sec/mm² isgenerally found for pixel size on the order of 1.0 mm and thicknesses oforder 3.0-5.0 mm. This limit is directly related to the chargecollection time for CZT. Higher flux rates are theoretically possibleusing of smaller pixels. Each pixel has a size-independent count ratelimit set by the charge collection time. The saturation flux rate is setby the count rate limit divided by the area of the pixel. Therefore, thesaturation flux rate increases as the pixel size decreases. Smallerpixels are also desirable because they make available higher spatialresolution information which can result in high resolution images.However, small pixel size results in higher cost and there are morechannels per unit area which need to be connected to readoutelectronics.

In addition, smaller pixels or detector elements have larger perimeterto area ratios resulting in more cross-talk. The perimeter is a regionwhere charge is shared between two or more pixels (i.e. cross-talk).This sharing of charge results in incomplete energy information and/or amiscount of x-ray photons because the readout electronics are notconfigured to combine simultaneous signals in neighboring pixels. Veryhigh flux rates are possible with thin, photon counting, directconversion silicon layers with pixel size <0.1 mm, but there is notsufficient stopping power in these thin layers to stop the x-rays. Forintegrating detectors, the size of the detector pixel and design of thepreamplifier are balanced to handle an x-ray flux rate expected duringimaging. For CT, the flux rate capability of the detector withintegrating electronics is generally of the order 10⁹ photons/sec/mm².For x-ray projection imagers operating with charge storage, integratingdetectors, the flux rate capability is only of the same order. Forphoton counting detectors using scintillators and one ofphotodiodes/APDs/photomultipliers, the dead time of the x-ray conversionlayer is very fast and the dead time is usually related to the bandwidthof the electronic readout, which can also be relatively high. Theproblem with these detectors is varied. In the case of photodiode, theelectronic gain is not sufficient to overcome the electronic noise. Inthe case of APDs, there is additional gain but with associatedgain-instability noise, temperature sensitivity and reliability issues.In the case of photomultiplier tubes, these devices are too large andcostly for high resolution detectors covering large areas.

Detector saturation can affect image quality by constraining the numberof photons used to reconstruct the image and introducing imageartifacts. A minimum image quality, therefore a minimum flux rate, isrequired to make use of the images. In this regard, when setting theconfiguration of the system such that sufficient flux is received at onearea of the detector, then it is likely that another area of thedetector will receive higher flux, and possibly, high enough to saturatethe detector in this area. Higher flux in these other areas is notnecessary for the image quality; however, the loss of data due todetector saturation may need to be addressed through correctionalgorithms in order to reduce image artifacts. For CT imaging, thereconstruction is not tolerant to missing or corrupted data. Forexample, if the center of the detector is illuminated with a minimumflux for image quality purposes, and if the illuminated object iscompact, then detector cells at and beyond the periphery of the object'sshadow can be saturated due to thin object thickness in these projecteddirections. The reconstruction of the data set with these uncorrectedsaturated values will cause severe artifacts in the image.

A number of imaging techniques have been developed to address saturationof any part of the detector. These techniques include maintenance of lowx-ray flux across the width of a detector array, for example, by usinglow tube current or current that is modulated per view. However, thissolution leads to increased scanned time. That is, there is a penaltythat the acquisition time for the image is increased in proportion tothe nominal flux needed to acquire a certain number of x-rays that meetimage quality requirements.

With respect to combined CT and SPECT or CT and PET imaging, theavailability of an energy discriminating detector with high flux ratecapability provides the opportunity for a shared detector. The x-rayphoton energies of SPECT are similar to those in CT, such that asemiconductor layer thickness can be designed to meet the requirementsof both CT and SPECT. However, for PET, the photon energies are at 511eV, about 5 times higher than that used for CT and SPECT.

It would therefore be desirable to design a direct conversion, energydiscriminating CT detector that does not saturate at the x-ray photonflux rates typically found in conventional CT systems. It would befurther desirable to design an x-ray management system that accommodatesvariations in x-ray flux across a CT detector assembly and compensatesfor over-ranging or saturating detectors. Such a detector and fluxmanagement system would allow the use of the same detector for both CTand SPECT imaging.

BRIEF DESCRIPTION OF THE INVENTION

The present invention is directed to a multilayer CT detector thatperforms at very high count rates that overcomes the aforementioneddrawbacks.

A CT detector capable of energy discrimination and direct conversion isdisclosed. Also, a dual-modality detector capable of both CT and singlephoton emission computed tomography (SPECT) detection is disclosed. Thedetector includes multiple layers of semiconductor material of varyingthicknesses throughout the detector. In this regard, the detector isconstructed to be segmented in the x-ray penetration direction tooptimize count rate performance as well as avoid saturation. Further,the CT detector is constructed such that data corresponding to saturatedregions can be estimated or otherwise determined from unsaturatedregions. Additionally, the CT detector may be fabricated so as to havemultiple detector elements or sub-pixels per contact area. In thisregard, a dynamic and flexible combining of the outputs of theindividual detector elements can be carried out to inhibit use of datafrom a saturated detector element.

The CT detector supports not only x-ray photon counting, but energymeasurement or tagging as well. As a result, the present inventionsupports the acquisition of both anatomical detail as well as tissuecharacterization information. In this regard, the energy discriminatoryinformation or data may be used to reduce the effects of beam hardeningand the like. Further, these detectors support the acquisition of tissuediscriminatory data and therefore provide diagnostic information that isindicative of disease or other pathologies. For example, detection ofcalcium in a plaque in a view is possible. This detector can also beused to detect, measure, and characterize materials that may be injectedinto a subject such as contrast agents and other specialized materialssuch as targeting agents. Contrast materials can, for example, includeiodine that is injected into the blood stream for better visualization.

Therefore, in accordance with one aspect of the present invention, a CTdetector is disclosed and includes a first direct conversion layerhaving a first array of electrical contacts and constructed to directlyconvert radiographic energy to electrical signals representative ofenergy sensitive radiographic data. The first direct conversion layer isalso designed to saturate at a first saturation threshold. The CTdetector further includes a second direct conversion layer having asecond array of electrical contacts and constructed to directly convertradiographic energy to electrical signals representative of energysensitive radiographic data. The second direct conversion layer isdesigned to saturate at a second saturation threshold different from thefirst saturation threshold.

In accordance with another aspect, the present invention includes aradiographic imaging system having a radiation source to projectradiographic energy toward a subject to be scanned and a detectorassembly to receive radiographic energy from the radiation source andattenuated by the subject. The detector assembly includes an array ofdetectors whereby each detector is designed to provide photon countand/or energy discriminatory output. The imaging system also includes acomputer programmed to detect pile-up and over-ranging in a section of agiven detector and determine appropriate output for the over-rangingsection of the given detector from non-over-ranging sections of thegiven detector.

According to another aspect, the present invention includes a CT scannerhaving a rotatable gantry with an opening to receive a subject to bescanned. The CT scanner also includes an x-ray source configured toproject x-rays toward the subject as well as a detector array having aplurality of detectors designed to provide energy sensitive output inresponse to detected x-rays. A data acquisition system (DAS) isconnected to the detector array and configured to receive the detectoroutputs. The CT scanner also includes an image reconstructor connectedto the DAS and configured to reconstruct an image of the subject fromthe detector outputs received by the DAS. The CT scanner furtherincludes means for determining an output for a given detector of thedetector array when a portion of the detector has reached an x-raysaturation state.

Various other features and advantages of the present invention will bemade apparent from the following detailed description and the drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

The drawings illustrate one preferred embodiment presently contemplatedfor carrying out the invention.

In the Drawings:

FIG. 1 is a pictorial view of a CT imaging system.

FIG. 2 is a block schematic diagram of the system illustrated in FIG. 1.

FIG. 3 is a perspective view of one embodiment of a CT system detectorassembly.

FIG. 4 is a perspective view of a CT detector incorporating the presentinvention.

FIG. 5 is illustrative of various configurations of the detector in FIG.4 in a four-slice mode.

FIG. 6 is a partial perspective view of a two-layer detector inaccordance with the present invention.

FIG. 7 is a cross-sectional view of FIG. 6 taken along lines 7-7thereof.

FIGS. 8-10 illustrate cross-sectional views of direct conversiondetectors in accordance with several additional embodiments of thepresent invention.

FIG. 11 is a cross-sectional view illustrating signal feedthroughs thatare created in another embodiment of the invention.

FIG. 12 is a cross-sectional schematic view of a CT detector inaccordance with another embodiment of the present invention.

FIG. 13 is a cross-sectional view of an alternate embodiment of aportion of a CT detector according to the present invention.

FIG. 14 is a cross-sectional view of a portion of a CT detector inaccordance with yet another embodiment of the present invention.

FIG. 15 is a cross-sectional view of yet another embodiment of a CTdetector according to the present invention.

FIG. 16 is a cross-sectional schematic view similar to FIGS. 13-15illustrating an alternate embodiment of the present invention.

FIG. 17 is a perspective view of a portion of a CT detector with itscomponents oriented in a vertical arrangement.

FIG. 18 is a top view schematic illustrating sub-pixelization of adetector element area according to the present invention.

FIG. 19 is a top view of a single CT detector element area illustratingasymmetrical sub-pixelization thereof in accordance with anotherembodiment of the present invention.

FIG. 20 is a top view of an alternate asymmetrical sub-pixelization fora single CT detector element area in accordance with an alternateembodiment of the present invention.

FIG. 21 is a block schematic illustrating combining the output of eachsub-pixel of a given CT detector element area in accordance with anotherembodiment of the present invention.

FIG. 22 is a circuit schematic illustrating a flexible binning ofsub-pixel outputs of a given CT detector element area in accordance withanother embodiment of the present invention.

FIGS. 23-24 are circuit schematics illustrating binning of sub-pixeloutputs of a given CT detector element area based on the saturationstate of each sub-pixel according to a further embodiment of theinvention.

FIG. 25 is a pictorial view of a CT system for use with a non-invasivepackage inspection system.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT

The operating environment of the present invention is described withrespect to a four-slice computed tomography (CT) system. However, itwill be appreciated by those skilled in the art that the presentinvention is equally applicable for use with single-slice or othermulti-slice configurations. Moreover, the present invention will bedescribed with respect to the detection and conversion of x-rays.However, one skilled in the art will further appreciate that the presentinvention is equally applicable for the detection and conversion ofother radiographic energy.

Referring to FIGS. 1 and 2, a computed tomography (CT) imaging system 10is shown as including a gantry 12 representative of a “third generation”CT scanner. Gantry 12 has an x-ray source 14 that projects a beam ofx-rays 16 toward a detector assembly 18 on the opposite side of thegantry 12. Detector assembly 18 is formed by a plurality of detectors 20which together sense the projected x-rays that pass through a medicalpatient 22. Each detector 20 produces an electrical signal thatrepresents not only the intensity of an impinging x-ray beam but is alsocapable of providing photon or x-ray count data, and hence theattenuated beam as it passes through the patient 22. During a scan toacquire x-ray projection data, gantry 12 and the components mountedthereon rotate about a center of rotation 24.

Rotation of gantry 12 and the operation of x-ray source 14 are governedby a control mechanism 26 of CT system 10. Control mechanism 26 includesan x-ray controller 28 that provides power and timing signals to anx-ray source 14 and a gantry motor controller 30 that controls therotational speed and position of gantry 12. A data acquisition system(DAS) 32 in control mechanism 26 review data from detectors 20 andconverts the data to digital signals for subsequent processing. An imagereconstructor 34 receives sampled and digitized x-ray data from DAS 32and performs high speed reconstruction. The reconstructed image isapplied as an input to a computer 36 which stores the image in a massstorage device 38.

Computer 36 also receives commands and scanning parameters from anoperator via console 40 that has a keyboard. An associated cathode raytube display 42 allows the operator to observe the reconstructed imageand other data from computer 36. The operator supplied commands andparameters are used by computer 36 to provide control signals andinformation to DAS 32, x-ray controller 28 and gantry motor controller30. In addition, computer 36 operates a table motor controller 44 whichcontrols a motorized table 46 to position patient 22 and gantry 12.Particularly, table 46 moves portions of patient 22 through a gantryopening 48.

As shown in FIGS. 3 and 4, detector assembly 18 includes a plurality ofdetectors 20, with each detector including a number of detector elements50 arranged in a cellular array. A collimator (not shown) is positionedto collimate x-rays 16 before such beams impinge upon the detectorassembly 18. In one embodiment, shown in FIG. 3, detector assembly 18includes 57 detectors 20, each detector 20 having an array size of16×16. As a result, assembly 18 has 16 rows and 912 columns (16×57detectors) which allows 16 simultaneous slices of data to be collectedwith each rotation of gantry 12.

Switch arrays 54 and 56, FIG. 4, are multi-dimensional semiconductorarrays coupled between cellular array 52 and DAS 32. Switch arrays 54and 56 include a plurality of field effect transistors (FET) (not shown)arranged as multi-dimensional array and are designed to combine theoutputs of multiple cells to minimize the number of data acquisitionchannels and associated cost. The FET array includes a number ofelectrical leads connected to each of the respective detector elements50 and a number of output leads electrically connected to DAS 32 via aflexible electrical interface 58. Particularly, about one-half ofdetector element outputs are electrically connected to switch 54 withthe other one-half of detector element outputs electrically connected toswitch 56. Each detector 20 is secured to a detector frame by mountingbrackets 61.

It is contemplated and recognized that for some applications, the countrate limitation of the FET arrays may make them less desirable. In thisregard, as will be described, each detection pixel or cell is connectedto a channel of electronics.

Switch arrays 54 and 56 further include a decoder (not shown) thatenables, disables, or combines detector element outputs in accordancewith a desired number of slices and slice resolutions for each slice.Decoder, in one embodiment, is a decoder chip or a FET controller asknown in the art. Decoder includes a plurality of output and controllines coupled to switch arrays 54 and 56 and DAS 32. In one embodimentdefined as a 16 slice mode, decoder enables switch arrays 54 and 56 sothat all rows of the detector assembly 18 are activated, resulting in 16simultaneous slices of data for processing by DAS 32. Of course, manyother slice combinations are possible. For example, decoder may alsoselect from other slice modes, including one, two, and four-slice modes.

As shown in FIG. 5, by transmitting the appropriate decoderinstructions, switch arrays 54 and 56 can be configured in thefour-slice mode so that the data is collected about line C_(L) from fourslices of one or more rows of detector assembly 18. Depending upon thespecific configuration of switch arrays 54 and 56, various combinationsof detectors 20 can be enabled, disabled, or combined so that the slicethickness may consist of one, two, three, or four rows of detectorelements 50. Additional examples include, a single slice mode includingone slice with slices ranging from 1.25 mm thick to 20.0 mm thick, and atwo slice mode including two slices with slices ranging from 1.25 mmthick to 10.0 mm thick. Additional modes beyond those described arecontemplated.

As described above, each detector 20 is designed to directly convertradiographic energy to electrical signals containing energydiscriminatory or photon count data. The present invention contemplatesa number of configurations for these detectors, its components, and themanner in which data is read out. Notwithstanding the distinctionsbetween each of these embodiments, each detector does share two commonfeatures. One of these features is the multilayer arrangement ofsemiconductor films or layers. In a preferred embodiment, eachsemiconductor film is fabricated from Cadmium Zinc Telluride (CZT).However, one skilled in the art will readily recognize that othermaterials capable of the direct conversion of radiographic energy may beused. The other common feature between the various embodiments is theuse of interstitial or intervening metallized films or layers separatingthe semi-conducting layers. As will be described, these metallizedlayers are used to apply a voltage across a semiconductor layer as wellas collect electrical signals from a semiconductor layer. As will alsobe described, detectors with such a design have improved saturationcharacteristics and photon count fidelity.

It is generally well known that the charge collection time of asemiconductor layer is inversely related to the maximum periodic countrate saturation threshold (MPR) of the layer. A thinner layer will havefaster collection of charges and higher MPR. However, the thinner layerwill stop a smaller fraction of the incident x-rays. The chargecollection time is approximately proportional to a dimension d, which isthe smaller of either the thickness of the detector or the pixel contactsize, whereas the radiographic energy deposition efficiency isexponentially increasing with thickness. The count rate performance fora CZT detector may be defined by:

${MPR} = {\frac{\mu_{e}E}{d}.}$

From this definition, assuming an equal contact size and thickness ofd=0.3 cm and an electric field E of 1000 V/cm, and with a μ_(e) of about1000 cm²/(V sec), a maximum periodic count rate of 3.0 megacounts/secmay be achieved. Since the arrival of x-rays is not periodic but random,significant saturation effects will occur at 10× lower average rate. Inother words, the count rate of a CZT semiconductor layer that is 3.0 mmthick may have a count rate performance in the range of 0.3-3.0megacounts/sec. However, as will be described, constructing a directconversion semiconductor detector with multiple layers with thecumulative thickness of a single thicker layer can improve count rateperformance.

Moreover, thinner conversion layers not only improve count rateperformance due to a reduction of charge collection time, but alsoprovide an improvement in charge collection efficiency thereby reducingpolarization, detector count and energy response fidelity. Thinnerconversion layers also reduce charge sharing between pixel elementsthereby improving energy discrimination performance and spatialresolution.

Improvement in flux rate performance through the segmenting of thedetector into multiple thin layers can be attributed to a number offactors. First, having multiple layers divides the total flux rate amongthe different layers. Each layer will experience only a fraction of thetotal flux. For example, incomplete x-ray attenuation of the firstlayer, which is thin (relative to the attenuation depth of x-rays), willinsure that saturation of this layer will be at a higher count rate thanthat of a thick layer that stops all the x-rays.

A second factor is that the thickness of the layers can be configured toinsure that if one layer does saturate, another layer is non-saturatedand gives valid data for that view. For example, if one of the layers isconstructed such that it stops only 5% of the x-rays, then it willsaturate at 20× the flux rate a thick layer designed to stop all thex-rays. A third factor is that charge collection time decreases as layerthickness and pixel size decrease. The charge collection time isapproximately proportional to either the thickness or pixel contactsize, whichever is smaller, divided by the mobility and electric fieldacross the layer. Smaller thicknesses and/or pixel size gives higherflux rate limit for that layer.

A fourth factor is that thinner layers also yield a reduction incross-talk. The impact of pixel size on cross-talk is approximatelygiven by the effective perimeter area over the total pixel area. Thatis, cross-talk is scaled by a factor 4W*aT/W² where W is the pixel pitchand aT is a charge spreading length proportional to the layer thickness.Therefore, cross-talk decreases as the layer thickness decreases. Thecompeting effects of flux rate saturation and cross-talk can be tradedoff by study of their impact on the detective quantum efficiency,(DQE(f)) an important figure of merit for x-ray imaging detectors. DQEfalls off as a function of count rate, less so for thinner layers. Thedesign methodology for optimization of the number of layers and theirthickness is predicated upon obtaining the greatest count rate beforewhich the DQE(f) has decreased below any point on a certain thresholdcurve.

A fifth factor is the reduction in polarization due to the moreefficient collection of electrons and holes. In a thinner layer, theelectrons and holes are able to travel a smaller distance before beingcollected; therefore, the electron and holes are less susceptible totrapping.

In addition to these five factors for improved count rate limit upon useof thin layers, the flux rate limit (i.e. count rate per unit area) isimproved by using smaller pixel size which is favored in thin layersbecause of reduced crosstalk.

Referring now to FIG. 6, a portion of a two-layered CZT or directconversion detector 20 a in accordance with one embodiment of thepresent invention is shown in perspective. Detector 20 a is defined by afirst semiconductor layer 62 and a second semiconductor layer 64. Duringthe fabrication process, each semiconductor layer 62, 64 is constructedto have a number of electronically pixilated structures or pixels todefine a number of detection elements or contacts 65. This electronicpixilation is accomplished by applying a 2D array 67, 69 of electricalcontacts 65 onto a layer 62, 64 of direct conversion material. Moreover,in a preferred embodiment, this pixilation is defined two-dimensionallyacross the width and length of each semiconductor layer 62, 64.

Detector 20 a includes a contiguous high voltage electrode 66, 68 forsemiconductor layers 62, 64, respectively. Each high voltage electrode66, 68 is connected to a power supply (not shown) and is designed topower a respective semiconductor layer during the x-ray or gamma raydetection process. One skilled in the art will appreciate that each highvoltage connection layer should be relatively thin so as to reduce thex-ray absorption characteristics of each connection layer and, in apreferred embodiment, is a few hundred angstroms thick. As will bedescribed in greater detail below, these high voltage electrodes may beaffixed to a semiconductor layer through a metallization process.

Referring now to FIG. 7, a cross-sectional view of FIG. 6 taken alongline 7-7 thereof illustrates the relative thickness of eachsemiconductor layer 62, 64 in one embodiment. In addition, for thisembodiment, the pixel pitch and contact size are similar in both layersand about equal to the smaller thickness. Similarly to the high voltageelectrode layers 66, 68, the 2D contact arrays 67, 69 should also beminimally absorbent of radiographic energy. Each array or signalcollection layer is designed to provide a mechanism for outputting theelectrical signals created by the semiconductor layers to a dataacquisition system or other system electronics. As explained in moredetail later, a mechanism for flexibly combining signals for pixels indifferent layers is provided. One skilled in the art will appreciatethat many (possibly several hundred) interconnects (not shown) are usedto connect all the contacts 65 with the CT system electronics. Further,the cumulative 3.0 mm thickness of the semiconductor conversion layer issuch that 99% of the x-rays are absorbed. The absorption is calculatedwith a physics model for an x-ray spectrum typical of a medical CTacquisition, 120.0 kvp spectrum filtered by 20.0 cm.

In addition, as shown in FIG. 7, the thickness of the semiconductorlayers 62, 64 is different from one another but the pixel pitch andcontact size is similar. The layers are arranged with a specific orderwith regard to the x-ray direction so as to leverage the exponentialabsorption characteristic. For FIG. 7, when the x-rays are incidentupward toward the bottom common contact array, more x-rays are depositedin semiconductor layer 62 than in semiconductor layer 64. For example,assuming that semiconductor layer 62 has a thickness of one millimeter(mm) and semiconductor layer 64 has a thickness of 2.0 mm, semiconductorlayer 62 is expected to absorb about 92% of the x-rays whereas thesecond semiconductor layer 64 is expected to absorb about 7% of thex-rays. The combined total absorption for the two layers replicates the99% efficiency of a 3.0 mm layer. One important benefit of theconstruction compared to 3 mm thick single layer, is a decrease inpolarization effects for these two thinner layers. This benefit itselfwill allow operation with a tenfold increase in flux rate in mostpractical applications.

In addition, by combining the count response from the two layers with aspecific self-correction algorithm, the segmented detector, detector 20a, may be constructed to provide a tenfold increase in count rateperformance relative to a single 3.0 mm thick layer of semiconductormaterial. Consider, for example, a CT detector, as described herein canbe constructed to have a first layer absorbing 92% of the incident x-rayflux and second layer absorbing 7%. As a result, the second layer willsaturate at a flux rate at 14× higher than a 3.0 mm thick layer. As theincident flux rate increases, the second layer will saturate orover-range at an x-ray flux more than the x-ray flux required tosaturate a 3.0 mm thick layer. This variability in saturationcharacteristics of multiple layers of a single CT detector allows forthe output of an over-ranged layer to be estimated by the effectivesignal in a non-over-ranged or non-saturated layer. In this regard, asaturation state of a given detector layer is detected and, as a result,signal for the saturated layer, or equivalent 3.0 mm thick layer, isempirically estimated from the output of the non-saturated layer in thedetector.

An example of this self-correctability algorithm is that at high countrate above which the first layer is saturated, only the count responsefrom a second layer weighted by its fractional absorption is assigned tothat projection for each pixel. At low count rate, a weighted sum of theresponse from both layers is assigned to the projection for each pixel.A more sophisticated algorithm may combine the signals for the twolayers with weighting inverse to their DQE such that as the statisticalerror in one layer's value grows with increasing count rate, then itsvalue is added to the combined sum with reduced weight.

It is contemplated that a CT detector assembly could be constructed suchthat each CT detector is constructed with such over-rangecorrectability. However, it is also contemplated that only thosedetectors in the detector assembly typically associated withover-ranging are constructed with this over-ranging self-correctability.For instance, the periphery detectors of a detector assembly typicallyencounter higher flux conditions than the more centrally disposeddetectors. In this regard, the peripheral detectors can be constructedwith over-ranging self-correctability whereas the more centrally locateddetectors are not. Further, layers with other detection mechanisms anddetector materials having high count rate capability, but poor countrate and/or energy response characteristics, can be used in certainparts of the detector to estimate the count rate and energy response ofthe saturated layer.

Additionally, it is contemplated that a given CT detector may have morethan two semiconductor layers. In this regard, the effective signaloutput of two or more non-saturated layers could be used to estimate theoutput of the saturated layers. For instance, a detector may beconstructed with a first layer that has a 35× effective response, thesecond layer having a 10× effective response, and a third layer with aneffective response equivalent to that of the first layer. In thisregard, the first and third layers would saturate at higher x-ray fluxlayers than the second layer. Accordingly, when the second layer hasover-ranged or saturated, the output of the first and third layers canbe used to compensate or effectively determine the output of theover-ranging second layer.

Referring now to FIG. 8, another contemplated design for a CZT or directconversion detector is shown. In this embodiment, detector 20 b alsoincludes a pair of semiconductor layers 74, 76. In contrast to thepreviously described embodiment, detector 20 b includes a single, commonsignal collection layer or 2D contact array 78. This single, yet commonarray 78 is designed to collect electrical signals from bothsemiconductor layers 74, 76 and output those electrical signals to a DASor other system electronics. In addition, detector 20 b includes a pairof high voltage electrodes 80, 82. Each high voltage electrodeeffectively operates as a cathode whereas the contacts 79 of the 2Darray 78 operate as an anode. In this regard, the voltage applied viahigh voltage connections 80, 82 creates a circuit through eachsemiconductor layer to the signal collection contacts array 78.

Yet another contemplated embodiment is illustrated in FIG. 9. As shownin this embodiment, detector 20 c includes four semiconductor layers 84,86, 88, and 90. Detector 20 c further includes two electricallyconductive lines or paths 92, 94 that are electrically connected to highvoltage electrodes 87, 89, 91 as well as collection contact arrays 93,95. Electrically conductive path 92 receives and translates electricalsignals from contact arrays 93, 95. In this regard, a single data outputis provided to the CT system electronics. Similar to a single signalcollection lead, a single high voltage connection 94 is used to powerthe four semiconductor layers 84-90 via electrodes 87, 89, 91. Detector20 c only requires a single high voltage connection.

Referring to FIG. 10, a monolithic embodiment of the present inventionis shown. Similar to the embodiment of FIG. 7, detector 20 d includesfour semiconductor layers 96-102. Each semiconductor layer 96-102 isconnected to a pair of electrically conductive layers. In this regard,one electrically conductive layer is used for application of a voltagewhereas the other electrically conductive layer is used for collectionof the electrical signals generated by the respective semiconductorlayers. To minimize the number of electrically conductive layers,detector 20 d utilizes an alternating electrically conductive layerarchitecture. That is, every other electrically conductive layer is usedfor high voltage connection with the other electrically conductivelayers used for signal collection. In this regard, electricallyconductive layers 104, 106, and 108 are used for application of arelatively high voltage whereas layers 110 and 112 include contacts forsignal collection. As such, high voltage collection layers 104 and 108are used to apply a voltage to semiconductor layers 96 and 102,respectively. High voltage connection layer 106 is used to apply avoltage to semiconductor layers 98 and 100.

As described above, in a preferred embodiment, each semiconductor layeris constructed from CZT material. One skilled in the art will appreciatethat there are a number of techniques that may be used to construct sucha semiconductor. For example, molecular beam epitaxy (MBE) is one methodthat may be used to grow each thin layer of CZT material. Screenprinting of CZT particles in polymetric binder is a potentially lowcost, low temperature method of forming layers on a flexible wiringsubstrate. One skilled in the art will appreciate that a number oftechniques may be used to metallize the semiconductor layers to providethe electrically conductive connections heretofore described.

Further, metallization may also be used to provide signal feedthroughsfor the collection contacts as illustrated in FIG. 11. As shown, asingle layer of semiconductor material 114 is sandwiched between anarray 116 of collection contacts and a high voltage electrode layer 118.Prior to metallization of the semiconductor layer 114 to form collectioncontact array 116 and high voltage electrode layer 118, holes 120 may beetched or otherwise formed in semiconductor 114. The holes 120 may thenbe metallized to provide a signal feed path 122 from a respectivecollection contact 124. The signal feedthroughs or conductive paths 122are constructed within holes 120 so as to not be in contact with thenear-contiguous high voltage electrode layer 118. In this regard, signalruns may extend vertically or in the x-ray reception directionthroughout the detector to a bus (not shown) designed to translate theelectrical signals emitted by the individual collection contacts 124 tothe CT system electronics. As a result, a stacked arrangement of aseries of thin stacked layers in the x-ray direction is formed.

Heretofore, the present invention has been described with respect to amultilayer CT detector designed with different layer thicknesses butsimilar dimension of the pixel size.

The present invention has been described with respect a multi-layer CTdetector incorporating direct conversion, semiconductor layers withvarying thickness to reduce the likelihood of such an energy sensitiveCT detector saturating or over-ranging at the x-ray flux rates typicallyencountered with conventional CT scans. As will be describedhereinafter, however, the present invention is also directed to anenergy sensitive, over-ranging resistant CT detector that utilizesvariability in the electrical contacts of a multi-layer CT detector toimprove saturation characteristics of the CT detector.

Referring now to FIG. 12, a side elevational, exploded view of a portionof a CT detector incorporating the present invention is shown. As shown,the detector 126 is formed on a substrate 128 that is secured to adetector frame (not shown) via fasteners (not shown). Substrate 128supports a pair of detector layers 130 and 132. Each detector layer 130,132 is composed of a radiation conversion component and a signalcollection component. Layers 130, 132, which in the illustratedembodiment, have different thicknesses, are stacked in the x-raydirection 134 and separated from one another by a flex layer 136. The CTdetector 126 also includes a high voltage bias wire 138 connected todetector layer 130 to bias the detector assembly.

As referenced above, detector 126 includes a pair of detector layers 130and 132. The detector layers may comprise scintillator and photodiodesconsistent with conventional CT detectors or fabricated from directionconversion semiconductor material, such as CZT, coupled to a number ofdetector elements or pixels. As illustrated in FIG. 12, detector layer130 differs from detector layer 132 in the thickness of respectiveconversion components 140, 142 and the number and size of the respectivedetector element arrays 144, 146. As illustrated, detector element array144 has half the pixel pitch, or four times the number of detectorelements 148 than detector element array 146 for an equivalent area ofthe detector. Additionally, the contact area of detector elements 148 isone fourth that of detector elements 150. As will be described, thisvariation in the detector element arrays within a single CT detectorgreatly enhances the saturation characteristics of the detector.

By varying the size of the detector elements within a given detector126, the charge collection time associated with each layer of thedetector is varied. That is, one skilled in the art will readilyappreciate that charge collection time decreases as the thickness of aconversion layer decreases and the size of the detector elementdecreases. That is, the charge collection time of a detector layer isapproximately proportional to the thickness of the conversion layer ordetector element size, whichever is smaller, divided by the mobility andelectric field across the detector layer. The count rate saturationthreshold will be larger for smaller pixel size. Furthermore, smallerpixel area implies a higher flux rate saturation threshold relative fora given count rate saturation threshold in proportion to the areareduction. As such, as conversion layer thickness and/or detectorelement size decreases, the flux rate limit for the correspondingdetector layer increases, thereby, improving the saturationcharacteristics for that layer of the CT detector. This improvement andvariability in saturation characteristics allows for a detector to beconstructed where some layers withstand higher x-ray flux levels andprovide inputs to a self-correctability algorithm.

In the embodiment illustrated in FIG. 12, the voltage bias wire or lead138 extends from substrate 128 to direct conversion layer 130. It iscontemplated, however, as shown in FIG. 13 that the high voltage biaswire 138 may be placed on flex layer 136. Flex layer 136 constitutes arouting layer and is used to connect the individual detector elements148 to the readout electronics, i.e. DAS and image reconstructor, of theCT scanner 10. In the embodiment illustrated in FIG. 13, the highvoltage wire 138 may be metallized on surface of flex layer 136 and thedetector elements 148 may be metallized on an opposite surface.

FIG. 14 illustrates another orientation of the detector layers 130, 132and their respective components relative to one another. As shown, thedifferences in thickness between direct conversion components 140 and142 may be varied from that shown in FIG. 12 to achieve differentabsorption and flux rate characteristics. For example, assuming thatdetector elements 148 have a pixel size of 0.25 mm and detector elements150 have a pixel size of 1.0 mm, detector layer 130, assumed to have athickness of 0.4 mm, will stop one-half of the x-rays 134 impingedthereon and conversion layer 142, assumed to have a thickness of 4.6 mm,will stop the other half of x-rays 134 not absorbed by layer 140. Theabsorption is calculated with a physics model for an x-ray spectrumtypical of a medical CT system, 140 kvp spectrum filtered by 3.0 cm.Relative to a single layer 5.0 mm thick, both detector layers 130, 132will each have a two-fold improvement in flux rate capability due toincomplete absorption in each layer individually. Further, layer 130 has1/16 the area of a 1.0 mm pixel resulting in 16× higher flux ratesaturation threshold versus a 1.0 mm pixel. Moreover, the combination ofreduced layer thickness and reduced detector element size results in a4× reduced charge collection time for layer 130 relative to a single 5.0mm detector layer with 1.0 mm pixel size. The flux rate improvements dueto each of these mechanisms are multiplicative. As a result, thiscombination of incomplete absorption (2×), smaller area (16×) anddecreased charge collection time (4×), the total flux rate saturationthreshold for detector layer 130 may be 128× higher relative to a 5.0 mmthick layer with 1.0 mm detector element pitch. Further, layer 130 willhave less polarization due to improved charge collection efficiency inthe thin layer.

Referring now to FIG. 15, relative to the embodiment of FIG. 14, theorder of the layered detector has been reversed. This reversal resultsin 99% of x-ray absorption in detector layer 132, which has a detectorelement pitch of 1.0 mm. As such, only 1% of the x-rays will be left forabsorption in detector layer 130, which has detector element pitch of0.25 mm. Accordingly, the absorbed flux rate fraction for detector layer130 is 100× less than that of a single layer detector 5.0 mm thick and apitch of 0.25 mm. Detector layer 132 achieves a 6× increase in flux ratecapabilities due to 4× faster charge collection time and 1/16 the pixelarea. In total, the multi-layer detector has a 6400× improvement in fluxrate performance compared to a single layer detector 5.0 mm thick and adetector element pitch of 1.0 mm.

Referring now to FIG. 16, it is contemplated that a detector 126 may beconstructed to have more than two detector layers. For example, detector126 may be designed to have three separate detector layers 130(a),130(b), and 132. For purposes of illustration, the detector of FIG. 16is oriented similar to that shown in FIG. 15 with the addition ofanother detector layer. As such, the detector illustrated in FIG. 16includes a relatively thick conversion layer 140 and two relativelythinner conversion layers 142(a), 142(b). Moreover, the detector elementpitch for detector layer 132 is 4× that of detector layers 130(a) and130(b). It is contemplated that a detector with the configurationillustrated in FIG. 16 will operate differently than the detectorsheretofore described.

Specifically, for the detector 126 of FIG. 16, at lower x-ray fluxlevels, none of the detector layers will saturate and the data from thesmaller detector elements 148(a) of array 144(a) and detector elements148(b) of array 144(b) will be combined to provide a single signal. Thatis, assuming that the detector elements 148(a) and 148(b) are one-fourththe pitch of detector elements 150, the count data for detector elements148(a) and 148(b) will be binned in a 4×4 manner so as to be equivalentto the pitch of detector elements 150 of detector layer 132. Forintermediate flux levels, detector layer 132 will saturate and countdata from detector layers 130(a) and 130(b) will only be used. The thirdlayer may be constructed to have a saturation threshold of 1000×compared to a 5.0 mm thick, single layer detector assembly having a 1.0mm detector element pitch.

Referring now to FIG. 17, it is contemplated that a vertical arrangementof the components of a CT detector may be used to also achieveimprovements in count rate performance. Detector 126 includes threedirect conversion layers that are uniform in their thickness and size.The direct conversion layers 152 are separated from one another by aflex layer 154 and an array of detector elements 156. With thisconstruction the thickness of the conversion layers sets detectorelement pitch in one direction and the spacing between detector elementsdefines the pitch in the other direction. Moreover, with thisarrangement, response of the detector elements is a function of detectorelement “height” on a particular conversion layer. For example, adetector element in row 158 of elements having a size of 0.7 mm collectscharge from approximately 1% of the x-rays impinged on the detectorwhereas detector element in row 159 of size 4.3 mm absorbs approximately99% of the x-rays impinges on the detector. The flux rate saturation forrow 158 of elements is therefore 100× greater than for a single 5.0 mmthick detector with 1.0 mm detector element pitch.

As referenced above, the present invention is directed to achievingimprovement in saturation characteristics of a CT detector and assemblyusing multiple direct conversion layers. The present invention is alsodirected to achieving improvement in saturation characteristics of theCT detector through reduction of detector element size. Each detectorelement of a CT detector is commonly referred to as a “pixel” and, assuch, in one embodiment, the present invention is directed to the“sub-pixilation” of a pixel area.

Referring now to FIG. 18, a single pixel area 160 (shown in phantom) ispixilated into four equally sized sub-pixels 162. In the illustratedexample, each sub-pixel 162 is connected to a dedicated readout lead164. Because flux at a pixel is proportional to its area, the combinedflux rate saturation threshold of the four separate sub-pixels 162 is 4×that which would be achieved by a single pixel 160 covering the area ofthe four sub-pixels 162. In addition, each sub-pixel 162 will have afaster charge collection time because of its reduction in size relativeto the layer thickness. Faster charge collection time is indicative of alarger saturation flux rate limit over and above the improvement incount rate performance achieved simply by a reduction in detectorelement size. It is noted that since each sub-pixel 162 is similarlysized, the sub-pixels will saturate at roughly the same x-ray fluxlevel.

On the other hand and referring to FIG. 19, the area achieved by asingle pixel 160 may be pixilated into sub-pixels that have differentflux rate characteristics. For example, as shown in FIG. 19, sub-pixel162(a) is significantly larger than sub-pixel 162(b). This asymmetry insub-pixel size yields a composite pixel area with different saturationthresholds within the composite pixel area. Specifically, assuming thatsub-pixel 162(a) is 20× larger than sub-pixel 162(b), then sub-pixel162(a) will saturate at an x-ray flux threshold 20× that of sub-pixel162(b).

It is contemplated that any number of orientations may be implemented toorientate sub-pixel 162(a) relative to sub-pixel 162(b). In thearrangement illustrated in FIG. 20, it is believed that the centerplacement of sub-pixel 162(b) improves cross-talk characteristicsbetween the sub-pixels. That is, in the illustrated arrangement,sub-pixel 162(b) is less likely to be affected by neighboring sub-pixelsand, as such, may be more immune from cross-talk from sub-pixel 162(a)when compared to the arrangement of FIG. 19.

It is recognized that flux rate is not uniform across a CT detector. Inthis regard, the present invention also includes an x-ray fluxmanagement system that detects and/or anticipates saturation of a givenportion of a CT detector such that appropriate corrective measures maybe taken. For example, it is well-known that the extremities of a CTdetector assembly often will receive more x-ray flux than the centerportions of the CT detector assembly due to subject and pre-subjectfilter attenuation profiles. As such, it is contemplated thatpost-acquisition logic may be used to only use the output ofnon-saturated channels for image reconstruction. In another embodiment,saturation of given portions of the CT detector assembly is anticipatedand, as a result, a binning scheme is established such that thoseportions of the CT detector expected not to saturate are electricallyconnected to the scanner's DAS and those portions expected to saturateare not. In yet another embodiment, connectivity of the detectorelements to the system DAS is determined on a per view basis during dataacquisition. That is, previous view data and other priori information isused to connect the detector elements to the DAS. This scheme provides adynamic, yet flexible binning of the detector elements during dataacquisition. In another embodiment, connectivity of the detectorelements to the DAS is controlled in real-time. In this regard,connectivity can be changed during the acquisition of data for a givenview such that connections are opened if high photon rate is detected.

Shown in FIG. 21 is a schematic of a group of detector elements orsub-pixels 164 for a given layer or array of detector elements with eachsub-pixel 164 electrically routed to a respective data system input 166.In the illustrated example, the group is comprised of four sub-pixels164 and, as such, four outputs 168 are provided to four data systeminputs 166. The four outputs of the 4×4 DAS channels 166 are input to atruth table circuit 170. The output 172 of the truth table circuit 170is a linear combination of the four inputs 174 to the truth tablecircuit 170 depending on whether any of the inputs are saturated.Although each of the sub-pixels illustrated in FIG. 21 is designed tosaturate at the same x-ray flux, it is possible given the contour of thesubject being imaged and pre-subject filtering for one sub-pixel of thegroup to saturate without saturation of a neighboring sub-pixel.

The table below sets out a truth table for combining the outputs of thefour sub-pixels. In the truth table, a value of “1” is indicative ofnon-saturation whereas a value of “0” is. As such, if none of thesub-pixels is saturated, a value of “1” will be input to the truth tablecircuit 170 for each of the sub-pixels. The truth table indicates thatin such a circumstance the outputs from all the sub-pixels areconsidered acceptable and combined to provide a single output for thatgroup of sub-pixels. On the other hand, if channel “A” or, moreprecisely, the sub-pixel associated with channel “A”, saturates, but theother sub-pixels have not, then the sum of the non-saturated channels isoutput by the truth table circuit and the data associated with thesaturated channel is ignored. For instance, assuming a pixel areacomposed of one sub-pixel that has a higher x-ray flux saturationthreshold higher than another sub-pixel within the pixel area, whenx-ray flux is low, both sub-pixels provide a valid output that is summedby the truth table circuit to provide a single output comprised of thecount data from both sub-pixels. When the x-ray flux reaches a level tosaturate only one of the sub-pixels, data from the non-saturatedsub-pixel is the only data output by the truth table circuit.

TABLE 1 LOGIC MAP A B C D Output 1 1 1 1 A + B + C + D 1 1 1 0 A + B + C1 1 0 1 A + B + D 1 1 0 0 A + B 1 0 1 1 A + C + D 1 0 1 0 A + C 1 0 0 1A + D 1 0 0 0 A 0 1 1 1 B + C + D 0 1 1 0 B + C 0 1 0 1 B + D 0 1 0 0 B0 0 1 1 C + D 0 0 1 0 C 0 0 0 1 D 0 0 0 0 Flag

It is recognized that a number of techniques may be used to determinesaturation of a given sub-pixel. For example, the count rate data for agiven sub-pixel may be compared to a threshold and if the count ratedetermined by the sub-pixel exceeds the threshold, a saturation value of“0” will be input to the truth table circuit for that sub-pixel. Forinstance, if the detector system is designed to count photons usingdirect conversion detectors with a one million count per secondsaturation threshold, then this threshold would be the threshold levelimposed on each sub-pixel, or some percentage thereof to provide amargin less than the saturation threshold.

It is also contemplated that a flexible binning of sub-pixels within agiven pixel area, such as that described above, may be achieved tofurther enhance the ability of the detector to output photon count datadespite saturation of some portions of the detector. That described withrespect to FIG. 21 was a signal management scheme that utilized a singleDAS channel for each sub-pixel. However, given the number of sub-pixelswithin a single CT detector, a single DAS channel per sub-pixel may notbe feasible. Accordingly, the present invention also contemplates asignal control scheme that utilizes one DAS channel for a group ofsub-pixels. In this regard, the number of DAS channels needed may beequivalent to that needed for a CT detector not incorporatingsub-pixilation.

Referring now to FIG. 22, a switch network-based signal managementsystem is shown whereupon more than one sub-pixel is dynamicallycontrolled to be connected to a DAS input channel. In this regard, theoutput of each sub-pixel 164 of a given group of sub-pixels is input toa switch network 176. The switch network is designed to reconnect theoutputs of the sub-pixels based on a saturation state of the sub-pixels.The switch network may utilize a truth table to dynamically controlconnectivity of the sub-pixel output. For those sub-pixels that havesaturated, the switch network will discard them such that onlynon-saturated data is included in output 180.

For example, at low x-ray flux, none of the sub-pixels will saturateand, as such, the output 168 from all the sub-pixels 164 will becombined into a single output 180 that is input to DAS 182. DAS 182includes a signal shaper 184 constructed to extract single photon eventsfrom the output of the switch network. It is recognized that alow-noise/high speed charge amplifier (not shown) may be connected toreceive the output of the switch network. The output of the amplifier isthen input to signal shaper 184. Signal shaper 184 provides an input toan energy level discriminator 186. Energy level discriminator 186 isconnected to signal shaper 184 and is designed to filter photons basedon their energy level relative to one or more thresholds. To this end,those photons having energy levels outside a desired range are excludedfrom counting and processing for image reconstruction. Minimally,discriminator 186 is designed to exclude those photons having an energylevel corresponding to noise in the system. It is contemplated thatmultiple thresholds may be used to define energy level ranges. Countingregister 188 receives those photons not filtered out by energy leveldiscriminator 186 and is constructed to count the number of photonsreceived at the detector and provide a corresponding output.

DAS 182 counts the number of photons for the given pixel area 165comprised of the given sub-pixels 164. Since the switch network will notconnect the output of a given sub-pixel if it has saturated, DAS willdetermine a photon count only from the non-saturated sub-pixels. Whileonly four sub-pixels are shown, it is contemplated that a given pixelarea may be sub-pixilated into less or more than four sub-pixels.

Two switch states are illustrated in FIGS. 23 and 24. As show in FIG.23, if sub-pixel identified with input “A” is not saturated and allother sub-pixels have saturated, then the output of sub-pixel “A” willonly be used for photon counting. As shown in FIG. 24, the switchnetwork may combine any combination of sub-pixel outputs, such assub-pixels “A” and “C” if those are the only non-saturated sub-pixels.

Referring now to FIG. 25, package/baggage inspection system 190incorporating a detector consistent with described herein is shown.System 190 includes a rotatable gantry 192 having an opening 194 thereinthrough which packages or pieces of baggage may pass. The rotatablegantry 192 houses a radiation source 196 as well as a detector assembly198. A conveyor system 200 is also provided and includes a conveyor belt202 supported by structure 204 to automatically and continuously passpackages or baggage pieces 206 through opening 194 to be scanned.Objects 206 are fed through opening 194 by conveyor belt 202, imagingdata is then acquired, and the conveyor belt 202 removes the packages206 from opening 194 in a controlled and continuous manner. As a result,postal inspectors, baggage handlers, and other security personnel maynon-invasively inspect the contents of packages 206 for explosives,knives, guns, contraband, etc.

An x-ray detector module is disclosed that includes a stacked sequenceof two or more x-ray detection layers, each containing an electricalcontact array. Interposed between the detection layers is an electricalrouting material for connecting contacts on the converter layer toreadout electronics. The number and thickness of the layers, and thepixel size on each layer is constructed to improve the count ratefidelity and resolution of the imaging system. In one embodiment, thetop (most proximate to the x-ray source) layer is thinner so as to stoponly a fraction of the incident x-rays. X-rays passing through the firstlayer are stopped in one of the subsequent layers. Each modulecontaining a sequence of layers is able to accommodate a higher fluxwithout saturation than an equivalent detector with one layer ofequivalent stopping power. The first layer, being thin and lesssusceptible to cross-talk, will also be capable of higher resolutionimaging with smaller pixels.

Additionally, it is disclosed that the signal from several sub-pixelsmaking up a composite pixel area may be weighted and combined in adynamic and real-time configurable manner into a single input channel ofa DAS. The configuration is chosen so as to include as many sub-pixelsinputs from a composite pixel unit, but to also insure that the DASchannel is not saturated and that the sub-pixels included are notsaturated or corrupted. This is particularly advantageous where thesub-pixels, which make up the composite pixel area, have dissimilarsaturation thresholds with respect to x-ray flux rate. For example, adetector element comprised of two sub-pixels, one with relatively largercoverage area and another with a relatively smaller coverage area, willhave differing saturation characteristics across the cumulative coveragearea due to area ratio and small pixel effect increase in the chargecollection time.

It is also disclosed that with a detector composed of superimposedlayers it is possible to combine signal from sub-pixels in one layerwith those of another layer but in a way that excludes saturated orotherwise corrupted sub-pixels. Four exemplary control configurationsare disclosed: (1) DAS-per-channel whereupon post-acquisition logic addsonly non-saturated channels to digital output; (2) Static, flexiblebinning, connectivity of sub-pixels to DAS established beforeacquisitions based on expected flux level based on a priori information;(3) Scan-Dynamic, flexible binning, sub-pixilation connectivity for eachview established during scan acquisition based on previous view data anda priori information; and (4) View-Dynamic, flexible binning,sub-pixilation connectively changed during a view starting with expectedbest case (e.g. starts with maximal connectivity and removes connectionsif high photon rate detected).

Therefore, the present invention includes a CT detector. The CT detectorincludes a first direct conversion layer having a first array ofelectrical contacts and constructed to directly convert radiographicenergy to electrical signals representative of energy sensitiveradiographic data. The first direct conversion layer is also designed tosaturate at a first saturation threshold. The CT detector furtherincludes a second direct conversion layer having a second array ofelectrical contacts and constructed to directly convert radiographicenergy to electrical signals representative of energy sensitiveradiographic data. The second direct conversion layer is designed tosaturate at a second saturation threshold different from the firstsaturation threshold.

The present invention also includes a radiographic imaging system havinga radiation source to project radiographic energy toward a subject to bescanned and a detector assembly to receive radiographic energy from theradiation source and attenuated by the subject. The detector assemblyincludes an array of detectors whereby each detector is designed toprovide photon count and/or energy discriminatory output. The imagingsystem also includes a computer programmed to detect over-ranging in asection of a given detector and determine output for the over-rangingsection of the given detector from non-over-ranging sections of thegiven detector.

The present invention further includes a CT scanner having a rotatablegantry with an opening to receive an object to be scanned. The CTscanner also includes an x-ray source configured to project x-raystoward the object as well as a detector array having a plurality ofdetectors designed to provide energy sensitive output in response todetected x-rays. A data acquisition system is connected to the detectorarray and configured to receive the detector outputs. The CT scannedalso includes an image reconstructor connected to the DAS and configuredto reconstructed image of the object from the detector outputs receivedby the DAS. The CT scanner further includes means for determining anoutput for a given detector of the detector array when a portion of thedetector has reached an x-ray saturation state.

The present invention has been described in terms of the preferredembodiment, and it is recognized that equivalents, alternatives, andmodifications, aside from those expressly stated, are possible andwithin the scope of the appending claims.

1. An imaging system comprising: a radiographic energy source configuredto emit a flux of x-rays that exceeds a first saturation threshold; adetector comprising: a stacked arrangement of a first direct conversionlayer and a second direct conversion layer; the first direct conversionlayer having a first array of electrical contacts and constructed todirectly convert radiographic energy to electrical signalsrepresentative of energy sensitive radiographic data, and having asaturation threshold; and the second direct conversion layer having asecond array of electrical contacts and constructed to directly convertradiographic energy to electrical signals representative of energysensitive radiographic data, and having a saturation threshold differentfrom the saturation threshold of the first direct conversion layer;wherein the first direct conversion layer is situated between theradiographic energy source and the second direct conversion layer; and acontroller configured to correct saturated data in one of the first andsecond direct conversion layers using unsaturated data acquired from theother of the first and second direct conversion layer.
 2. The imagingsystem of claim 1 wherein the detector further comprises an electricalreadout layer electrically connected to the first and the second arrayof electrical contacts and designed to route electrical signals from thecontact arrays to a data acquisition system.
 3. The imaging system ofclaim 1 wherein the first direct conversion layer is constructed to haveless x-ray stopping power than the second direct conversion layer. 4.The imaging system of claim 1 wherein the first array of electricalcontacts has more electrical contacts than the second array ofelectrical contacts.
 5. The imaging system of claim 1 wherein thedetector has more than two direct conversion layers.
 6. The imagingsystem of claim 1 wherein the second direct conversion layer isconstructed to over-range before the first direct conversion layer. 7.The imaging system of claim 6 wherein the first direct conversion layeris further designed to output electrical signals of which at least aportion is used to estimate electrical signal output of the seconddirect conversion layer when the second direct conversion layer is in anover-ranged state.
 8. The imaging system of claim 1 wherein eachelectrical contact includes multiple contact elements.
 9. The imagingsystem of claim 8 wherein each contact element is constructed to providean output that is combinable with an output of one or more other contactelements to a common electrical contact.
 10. The imaging system of claim1 configured to count photons received and associate an energy level toeach photon counted.
 11. The imaging system of claim 1 wherein theimaging system is a CT imaging system.
 12. A radiographic imaging systemcomprising: a radiation source to project radiographic energy toward asubject to be scanned; a detector assembly having a first layer and asecond layer and configured to receive radiographic energy from theradiation source and attenuated by the subject, each of the first andsecond layers having an array of detectors designed to output photoncount and energy level per photon; and a computer programmed to: detectan over-ranging section in the first layer; and determine output for theover-ranging section of the first layer from one or morenon-over-ranging sections of the second layer.
 13. The system of claim12 wherein each of the first and second layers comprises semiconductorsdesigned to directly convert received photon events of radiographicenergy to individually tagged electrical signals tagged with energylevel data.
 14. The system of claim 13 wherein the first layer has athickness different from a thickness of the second layer.
 15. The systemof claim 13 wherein each semiconductor layer includes a number ofcontacts, each contact designed to provide photon count and/or energylevel output.
 16. The system of claim 15 wherein each contact isdesigned to saturate at a given level that varies among the contacts ofa respective semiconductor layer.
 17. The system of claim 15 whereineach contact comprises multiple contact elements, each contact elementdesigned to provide photon count and/or energy level output.
 18. Thesystem of claim 17 wherein the computer is further programmed to combinethe outputs of more than one contact element to provide a single outputfor a respective contact.
 19. The system of claim 18 wherein thecomputer is further programmed to determine if a given contact elementhas become saturated and, if so, not incorporate the output of thesaturated contact element with the output of the respective contact. 20.A method of CT imaging comprising: receiving x-rays at a first layer ofa CT detector that emit from an x-ray source; receiving x-rays at asecond layer of a CT detector that emit from the x-ray source and passthrough the first layer; detecting an over-ranging detector in the firstlayer; correcting an output of the over-ranging detector from at leastone non-over-ranging detector of the second layer; and displaying animage reconstructed using the corrected output.